The present embodiments relate to an improved antenna circuit for a magnetic resonance imaging (MRI) system that controls an antenna assembly or reads-in signals of the antenna assembly.
MRI systems from the prior art include a single transmitter that is provided to generate a substantially homogeneous high frequency field for applying magnetic resonance. The associated transmitting antenna (e.g., a Body Coil (BC)) is permanently installed in a magnet and in gradient coils. To receive magnetic resonance signals, a multichannel assembly of receiving antenna (e.g., local coils) closely applied to the patient is used by contrast. The antennae that are not used in each case may be switched off or de-tuned by a PIN diode switch.
Applying the local coils to the object to be examined (e.g., to the patient) and forwarding the receiving signals to the patient couch is undesirable due to the complicated cable routing. A permanently installed receiving antenna assembly including very low-noise antenna elements using a “Remote Body Array” has, therefore, been proposed. A remote body array is described by way of example in US 2010/0213939 A1. A radial installation space for a cylindrical “Remote Body Array” (RBA) is radially inwardly limited because an optimally large patient opening is desirable. Larger diameters of the magnet or gradient systems lead to greatly increasing costs. The receiving elements of the RBA are therefore to be placed close to the transmitting antenna (BC). This results in a strong magnetic coupling between the transmitting coil (BC) and the RBA, whereby the requirements made on the de-tuning switch in the transmitting antenna (BC) (e.g., on the reduction of losses) are increased. Eddy currents may form on conductor structures of the transmitting antenna (BC). The eddy currents generate loss resistances even without participation of the switch and thereby couple noise into the receivers of the RBA. These losses are a problem because the requirements on the quality of all RBA elements are significantly higher, owing to the large spacing from the patient, than with local coils applied to the patient.
It has been proposed that the transmitting antenna be dispensed with, and the receiving elements of the RBA also be used for transmission. This solution has the advantage that the RBA may be made thicker and therefore more efficient since the additional space for the transmitting antenna (BC) may be omitted. Each antenna element receives a transceiver switch. Since this changeover switch does not have to be located in a high-quality resonance current circuit, but may be arranged after the transformation (e.g., at an impedance of 50 Ohm), a loss contribution of the changeover switch is no higher than is the case with conventional antenna designs. Receiving branches of the switches are forwarded to one receiver per element.
FIG. 11 shows an assembly from the prior art. An antenna element 102 is connected to a transceiver switch 104. An output of the transceiver switch 104 is connected to an input of an amplifier 110. The input of the transceiver switch 104 is connected to a distribution network, to which a plurality of antennae 102 are connected by corresponding transceiver switches 104. An output of a transmitting amplifier 108 is connected to an input of the distribution network 106.
Known solutions to this distribution are cascades including power splitters and phase shifters (e.g., Wilkinson splitters and Butler matrix) that may be inserted in lines with an impedance of 50 Ohm.
Such supply networks have a well-defined output impedance (e.g., 50 Ohm), with which each of the antennae is supplied. Each antenna may be optimally adapted to this resistance with respect to power. Different load resistances may occur at the individual antennae 102, however, due to different patient sizes and patient positions. A significant mutual coupling of the antennae 102 is to be anticipated. Even with a distribution of the waves converging on the antennae that is homogenous per se, very uneven current distributions may form. The uneven current distributions may lead to an inhomogeneous distribution of a spin-excitation and a patient power loss (SAR). This problem does not occur with the conventional high-pass or low-pass birdcage antenna assemblies used for transmission, since the resonance frequencies of the higher spatial frequency modes are very distant. The elements of a transceiver assembly are to be decoupled from each other to achieve the functionality of the assembly, so different current distributions may form.
There is therefore a requirement for a distribution network that, when transmitting signals using an antenna assembly, enforces a fixed and load-independent relationship of the currents in the antenna elements with respect to each other.
An equal current may be attained in a plurality of coils, for example, by a series connection. Even if each coil is to be supplemented by a capacitor to form a series resonance circuit, the resonance circuits may be connected in series. However, with high frequencies, a series connection may not be practicable owing to the distributed capacitors in the connecting lines.